The present embodiments relate to a high-frequency shield system for a magnetic resonance device.
In a magnetic resonance device, a body to be examined using a basic field magnet system may be exposed to a relatively high basic field magnetic field of 3 or 7 tesla, for example. In addition, a magnetic field gradient is created using a gradient system. Via a high-frequency transmission system, high-frequency excitation signals (HF signals) are transmitted using suitable antenna devices that may result in the nuclear spins of particular atoms resonantly excited by this high-frequency field being tilted by a defined flip angle with respect to the magnetic field lines of the basic magnetic field. This high-frequency excitation or the resulting flip angle distribution is also called nuclear magnetization. During the relaxation of the nuclear spin, high-frequency signals (e.g., magnetic resonance signals) are emitted. The high-frequency signals are received and further processed by suitable receiving antennas. The desired image data may be reconstructed from the raw data acquired in this way. The transmission of the high-frequency signals for nuclear spin magnetization may be effected using a “bodycoil.” A typical design for the body coil is a birdcage antenna that includes several transmission rods that are arranged running around a patient space (in which a patient is located during the examination) in the tomographic system parallel to the longitudinal axis. The antenna rods are each capacitively connected to one another in a ring shape on the end face. Apart from transmitting, this antenna may also be used to receive magnetic resonance signals.
Local coils may be used to receive the magnetic resonance signals and are applied directly to the body of the patient. The local coils may consist of a group of conductor loops (e.g., an antenna array), with the antenna conductor loops being individually operable. These antennas are designed with respect to antenna elements of the antennas such that the antenna elements may sensitively receive even low signals. The received low signals may be amplified and used as raw data. An antenna array such as this may form a relatively large surface antenna on the body of the examination object or patient.
By mounting the local coils close to the body, a maximized signal-to-noise ratio (SNR) may be obtained in the received signal and thus in the diagnostic information. Another advantage of an antenna array such as this with several individually operable conductor loops is that image acquisition is made significantly faster in the context of parallel imaging procedures and thus the exposure experienced by the patient may be reduced. By using spatially adjacent individual antenna elements, additional local information that complements the local resolution achieved by the gradient fields may be obtained. If the local coils are appropriately designed and wired, the local coils may also be used to transmit.
Mounting local coils close to the body also has certain practical disadvantages. Firstly, compared with other imaging modalities such as, for example, computed tomography systems, a great deal of additional time is needed when mounting the local coils. As a result, the magnetic resonance devices are busy for longer and are not available for other examinations. In addition, this extra waiting time may at least impose psychological stress on the patient. Secondly, mounting local coils on the body is uncomfortable and restrictive for the patient and, in extreme cases, may make examinations not only more protracted but even impossible. Thirdly, the local coils are wired in the patient table to the receiving devices of the magnetic resonance device. This wiring generates considerably higher costs when manufacturing a magnetic resonance device. The plugs and cable at the local coils are susceptible to abrasion. Because of these disadvantages, a technical solution for replacing local coils with antennas located further away from the body that may be mounted on the system side in the magnetic resonance tomography device rather than directly on the patient is desired.
Such a remote-body array (RBA) of individual receiving antennas is described, for example, in U.S. Ser. No. 12/392,537. It has been shown that as a result of the relatively large distance of such an RBA from the patient's body, both the induced MR received signal and the noise received from the patient's body is very small. Hence, the overall noise in the remote-body receiving coils is dominated by the thermal inherent noise of the loss resistors in the local coils. To be able to achieve a better signal-to-noise ratio, the inherent noise of the receiving antennas is to be reduced, either by reducing loss resistors of the receiving antennas or a temperature of the receiving antennas. U.S. Ser. No. 12/392,537 proposes to implement an RBA with extremely low inherent noise by cooling (e.g., with helium or nitrogen) or by using superconductors (e.g., high-temperature).
Also, the mirror currents that develop on a high-frequency shield that may be arranged between a transmitting or receiving antenna and the gradient coil mounted radially further outward, is a problem. This high-frequency shield is provided in order to shield the gradient coil from the high-frequency fields of the antenna arrangement. Such high-frequency shields between the bodycoil (e.g., the birdcage antenna) and the gradient coil system are also used in the magnetic resonance devices without RBA, which are standard in the art. These high-frequency shields may be designed in the form of slotted metal shield surfaces with two slotted shield surfaces being located in each case on different sides of a flexible printed circuit, and the slots in each case being arranged such that the slots are offset with respect to one another. The slots of one of the two shield surfaces are covered by the metal surface of the other shield surface, and vice versa. The slots may be bridged with capacitors that are conductive for high frequency but largely have a blocking effect for the frequencies of the gradient system. This provides that the high-frequency fields of the antenna are shielded from the gradient coil as desired, but the shield is transparent for the low-frequency gradient fields. These shield surfaces have an unavoidable surface resistance that in part causes feedback of the mirror currents on the HF shield into the antenna elements and thus when receiving, contributes to additional thermal noise in the antenna elements. This problem, by nature, becomes all the more critical if not only one transmitting antenna is used in the vicinity of the high-frequency shield, but an RBA with a plurality of receiving antenna elements is to be integrated there, and these antenna elements are to be extremely low-loss.
To reduce the losses in the high-frequency shield, the distance between the antenna arrangement and the high-frequency shield may be increased. However, because of the radial space present within the available inner chamber of a modern magnetic resonance scanner, this approach is of limited help, since either the diameter of the measurement chamber that is usable for the patient becomes smaller, which dramatically reduces patient comfort, or the basic field magnet is larger in size, resulting not only in higher costs, but also in problems when setting up the devices.